Transesophageal Ultrasound Using a Narrow Probe

ABSTRACT

Transesophageal echocardiography is implemented using a miniature transversely oriented transducer that is preferably small enough to fit in a 7.5 mm diameter probe, and most preferably small enough to fit in a 5 mm diameter probe. Signal processing techniques improve the depth of penetration to the point where the complete trans-gastric short axis view of the left ventricle can be obtained, despite the fact that the transducer is so small. The reduced diameter of the probe (as compared to prior art probes) reduces risks to patients, reduces or eliminates the need for anesthesia, and permits long term direct-visualization monitoring of patients&#39; cardiac function.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a continuation of application Ser. No. 12/732,937,filed Mar. 26, 2010, which is a divisional of application Ser. No.10/997,059, filed Nov. 24, 2004 (now U.S. Pat. No. 7,717,850), whichclaims priority to U.S. provisional application No. 60/525,330, filedNov. 26, 2003.

BACKGROUND

In the medical field, monitoring heart function impacts criticaldecisions that relate to patient care. One type of prior art heartmonitor is the intravascular/intracardiac ultrasound transducer (such asthe Accunav™ transducer). This type of transducer, however, is not wellsuited for transesophageal echocardiography because the transducerelements are oriented longitudinally instead of transversely, whichlimits the types of images that can be obtained. A second type of priorart heart monitor is the transesophageal echocardiography (TEE)transducer, which is transversely oriented. However, in order to producerepeatedly usable images, the azimuthal aperture of these transducersmust be quite large (e.g., 10-15 mm in diameter for adults), whichrequires a correspondingly large probe. Because of this large probe,conventional TEE often requires anesthesia, can significantly threatenthe airway, and is not well suited for long-term monitoring of theheart.

SUMMARY OF THE INVENTION

Transesophageal ultrasound imaging is implemented using a miniaturetransversely oriented transducer that is preferably small enough to fitin a 7.5 mm diameter probe, and most preferably small enough to fit in a5 mm diameter probe. Signal processing techniques provide improved depthof penetration, despite the fact that the transducer is so small.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is an overall block diagram of a system for monitoring cardiacfunction by direct visualization of the heart.

FIG. 2 is a more detailed view of the probe shown in the FIG. 1embodiment.

FIG. 3 is a schematic representation of a displayed image of thetrans-gastric short axis view (TGSAV) of the left ventricle.

FIG. 4 depicts the positioning of the transducer, with respect to theheart, to obtain the TGSAV.

FIG. 5 shows a plane that slices through the trans-gastric short axis ofthe heart.

FIG. 6A shows an optional probe interface configuration.

FIG. 6B is a graph of gain characteristics for a TGC amplifier.

FIGS. 7A, 7B, and 7C show a first preferred transducer configuration.

FIGS. 8A and 8B show a second preferred transducer configuration.

FIG. 9 shows the components of spatial resolution.

FIG. 10 shows the interaction between the shape of the resolution voxeland the boundary.

FIG. 11 shows the sector width.

FIG. 12 is a schematic illustration of the paths of the ultrasound beamas it is swept through the sector.

FIG. 13 is a schematic illustration of the samples that correspond to asection of one of the beams of FIG. 12.

FIG. 14 is a flowchart of a processing algorithm that uses frequencycharacteristics of the return signal.

FIG. 15 is a graph of a function that maps a gain factor onto an energyratio.

FIGS. 16A and 16B show two alternative transducer designs.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

FIG. 1 is an overall block diagram of a system that may be used forcontinuous long term monitoring of cardiac function by directvisualization of the heart. An ultrasound system 200 is used to monitorthe heart 110 of the patient 100 by sending driving signals into a probe50 and processing the return signals received from the probe intoimages, using the image processing algorithms described below. Theimages generated by those algorithms are then displayed on a monitor210, in any conventional manner.

FIG. 2 shows more details of the probe 50, which is connected to theultrasound system 200. At the distal end of the probe 50 there is ahousing 60, and the ultrasound transducer 10 is located in the distalend 64 of the housing 60. The next portion is the flexible shaft 62,which is positioned between the distal end 64 and the handle 56. Thisshaft 62 should be flexible enough so that the distal end 64 can bepositioned past the relevant anatomical structures to the desiredlocation, and the handle 56 facilitates the positioning of the distalend 64 by the operator. Optionally, the handle 56 may contain atriggering mechanism 58 which the operator uses to bend the end of thehousing 60 to a desired anatomical position as described below.

At the other end of the handle 56 is a cable 54, which terminates, atthe proximal end of the probe 50, at connector 52. This connector 52 isused to connect the probe 50 to the ultrasound system 200 so that theultrasound system 200 can operate the probe. Signals for the ultrasoundsystem 200 that drive the transducer 10 travel through the probe 50 viaappropriate wiring and any intermediate circuitry (not shown) to drivethe transducer 10, and return signals from the transducer 10 similarlytravel back through the probe 50 to the ultrasound system 200 where theyare ultimately processed into images. The images are then displayed onthe monitor 210 in a manner well known to persons skilled in therelevant art.

In the preferred embodiments, the housing 60 has an outer diameter ofless than 7.5 mm. The probe contains the ultrasound transducer 10 andconnecting wires, and the housing 60 can be passed through the mouth ornose into the esophagus and stomach.

The returned ultrasound signals are processed in the ultrasound system200 to generate an image of the heart. Preferably, additional signalprocessing is used to significantly improve image production, asdescribed below. FIG. 3 shows a displayed image of the trans-gastricshort axis view (TGSAV) of the left ventricle (LV), which is a preferredview that can be imaged using the preferred embodiments. The illustratedimage of the TGSAV appears in a sector format, and it includes themyocardium 120 of the LV which surrounds a region of blood 130 withinthe LV. The image may be viewed in real time or recorded for laterreview, analysis, and comparison. Optionally, quantitative analyses ofcardiac function may be implemented, including but not limited tochamber and vessel dimensions and volumes, chamber function, blood flow,filling, valvular structure and function, and pericardial pathology.

Unlike conventional TEE systems, the relatively narrow housing used inthe preferred embodiments makes it possible to leave the probe inposition in the patient for prolonged periods of time.

As best seen in FIGS. 4 and 5, the probe 50 is used to introduce andposition the transducer 10 into a desired location within the patient'sbody. The orientation of the heart within the chest cavity is such thatthe apex of the left ventricle is positioned downward and to the left.This orientation results in the inferior (bottom) wall of the leftventricle being positioned just above the left hemidiaphragm, which isjust above the fundus of the stomach. During operation, the transducer10 emits a fan-shaped beam 90. Thus, positioning the transducer 10 inthe fundus of the stomach with the fan-shaped beam 90 aimed through theleft ventricle up at the heart can provide a trans-gastric short axisview image of the heart 110. The plane of the fan-shaped beam 90 definesthe image plane 95 shown in FIG. 5. That view is particularly useful formonitoring the operation of the heart because it enables medicalpersonnel to directly visualize the left ventricle, the main pumpingchamber of the heart. Note that in FIGS. 4 and 5, AO represents theAorta, IVC represents the Inferior Vena Cava, SVC represents theSuperior Vena Cava, PA represents the pulmonary artery, and LVrepresents the left ventricle.

Other transducer positions may also be used to obtain different views ofthe heart, typically ranging from the mid-esophagus down to the stomach,allowing the operator to directly visualize most of the relevant cardiacanatomy. For example, the transducer 10 may be positioned in the loweresophagus, so as to obtain the conventional four chamber view.Transducer positioning in the esophagus would typically be done withoutfully flexing the probe tip, prior to advancing further into thestomach. Within the esophagus, desired views of the heart may beobtained by having the operator use a combination of some or all of thefollowing motions with respect to the probe: advance, withdraw, rotateand slight flex.

For use in adults, the outer diameter of the housing 60 is preferablyless than about 7.5 mm, more preferably less than about 6 mm, and ismost preferably about 5 mm. This is significantly smaller thanconventional TEE probes. This size reduction may reduce or eliminate theneed for anesthesia, and may help expand the use of TEE for cardiacmonitoring beyond its previous specialized, short-term settings. When a5 mm housing is used, the housing is narrow enough to pass through thenose of the patient, which advantageously eliminates the danger that thepatient will accidentally bite through the probe. Alternatively, it maybe passed through the mouth like conventional TEE probes. Note that the5 mm diameter of the housing is similar, for example, to typical NG(naso-gastric) tubes that are currently successfully used long-termwithout anesthesia in the same anatomical location. It should thereforebe possible to leave the probe in place for an hour, two hours, or evensix hours or more.

The housing wall is preferably made of the same materials that are usedfor conventional TEE probe walls, and can therefore withstand gastricsecretions. The wiring in the probe that connects the transducer to therest of the system may be similar to that of conventional TEE probes(adjusted, of course, for the number of elements). The housing ispreferably steerable so that it can be inserted in a relatively straightposition, and subsequently bent into the proper position after it entersthe stomach. The probe tip may be deflected by various mechanismsincluding but not limited to steering or pull wires. In alternativeembodiments, the probe may use an intrinsic deflecting mechanism such asa preformed element including but not limited to pre-shaped materials.Optionally, the probe (including the transducer housed therein) may bedisposable.

For imaging the TGSAV of the LV, the probe tip is preferably ultimately“ante-flexed” (flexed towards the front of the patient) approx. 70-110degrees. This may be implemented, for example, by building a triggerableante-flex (e.g., on the order of 70 degrees) into the probe through acombination of a pre-formed element, a device to prevent flexing duringinsertion and a trigger to release the preformed element from theinsertion limit once the probe is in the desired anatomic location.Optionally, a pull-wire may be used for steering to provide theadditional 0-40 degrees of flex after the transducer is lowered to theappropriate depth. The triggerable ante-flex component is preferablydesigned so that it will present little resistance to returning to theunflexed position during probe removal.

FIG. 6A shows an optional configuration that is similar to the FIG. 1embodiment, except that the circuitry that interfaces with the probe 50is relocated to an interface box 203. The rest of the ultrasound systemremains in the main processing unit 201, which communicates with theinterface box 203 via an appropriate cable 205. The interface box 203contains circuitry to amplify the signals from the transducer 10 and/ordigitize those signals. Using such an interface box advantageouslyprovides shorter signal paths for those parts of the circuit that aremost sensitive to electrical noise (i.e., where the signals are small).The transmit signals that drive the transducer 10 may also be generatedwithin the interface box 203 if desired.

Electrical noise may be further reduced using a variety of techniques.For example, in one embodiment, the interface box 203 houses apreamplifier that serves as the first stage in theamplification/processing chain, and separate power supplies are used forthe interface box and the main processing unit 201 to reduce electronicnoise pass-through. In another embodiment, the interface box 203 housesa preamplifier that serves as the first stage in theamplification/processing chain, and the preamplifier operates on batterypower. For both of these embodiments, time gain compensation (TGC) ispreferably implemented in that preamplifier. TGC compensates for thefact that the return signals from distant scatterers are weaker thanthose for nearby scatterers by increasing the gain for signals withlonger travel time. TGC may be implemented using conventional techniquesthat are well known to persons skilled in the relevant art. An exampleof suitable gain vs. delay characteristics for TGC is shown in FIG. 6B,where the x-axis represents the delay between transmission of theultrasound pulse and detection of the return signal, and corresponds todepth as follows:

Depth(in cm)=0.077 cm/μs×delay(in μs).

Implementing TGC in the preamplifier facilitates efficient digitization.The preamplifier may also provide amplitude compandoring (a form ofcompression) to further facilitate efficient digitization. Optionally,the preamplifier's output may be digitized in the interface box, inwhich case only digital signals would be sent from the interface box tothe main processing unit to further reduce electrical noise. Thesedigital signals may even be opto-isolated to eliminate all possibleelectrical connections in the return path, to reduce electrical noisepass-through further still.

The preferred embodiments described herein provide a good quality imageof the TGSAV of the LV from a transducer that is small enough to fit inthe narrow housing described above. FIGS. 7A-7C depict a first preferredtransducer 10. FIG. 7A shows the location of the transducer 10 in thedistal end of the housing 60, and also includes a top view 22 of thetransducer 10 surrounded by the wall of the housing 60 and a frontcutaway view 24 of the transducer 10.

As best seen in FIG. 7B, the azimuth axis (Y axis) is horizontal, theelevation axis (Z axis) is vertical, and the X axis projects out of thepage towards the reader. When steered straight forward by energizing theappropriate elements in the transducer, the beam will go straight outalong the X axis. The steering signals can also send the beam out atangles with respect to the X axis, in a manner well know to personsskilled in the relevant arts.

The transducer 10 is preferably a phased array transducer made of astack of N piezo elements L₁ . . . L_(N), an acoustic backing 12, and amatching layer in the front (not shown), in a manner well known to thoseskilled in the relevant art. As understood by persons skilled in therelevant arts, the elements of phased array transducers can preferablybe driven individually and independently, without generating excessivevibration in nearby elements due to acoustic or electrical coupling. Inaddition, the performance of each element is preferably as uniform aspossible to help form a more homogeneous beam. Optionally, apodizationmay be incorporated into the transducer (i.e., tapering the powerdriving transducer elements from a maximum at the middle to a minimumnear the ends in the azimuthal direction, and similarly tapering thereceive gain).

The preferred transducers use the same basic operating principles asconventional TEE transducers to transmit a beam of acoustic energy intothe patient and to receive the return signal. However, while the firstpreferred transducer 10 shown in FIGS. 7A-7C shares many characteristicswith conventional TEE transducers, the first preferred transducer 10differs from conventional transducers in the following ways:

TABLE 1 conventional TEE first preferred Feature transducer transducerSize in the transverse (azimuthal) 10-15 mm about 4-5 mm directionNumber of elements 64 about 32-40 Size in the elevation direction 2 mmabout 4-5 mm Front face aspect ratio about 1:5 about 1:1(elevation:transverse) Operating frequency 5 MHz about 6-7.2 MHzIn FIG. 7A, the elevation is labeled E and the transverse aperture islabeled A on the front cutaway view 24 of the transducer 10. Thelocation of the wall of the housing 60 with respect to the transducer 10can be seen in the top view 22.

FIG. 7C shows more details of the first preferred transducer 10. Notethat although only eight elements are shown in all the figures, thepreferred transducer actually has between about 32-40 elements, spacedat a pitch P on the order of 130 μm. Two particularly preferred pitchesare approximately 125 μm (which is convenient for manufacturingpurposes) and approximately 128 μm (0.6 wavelength at 7.2 MHz). When32-40 elements are spaced at a 125 μm pitch, the resulting azimuthaperture A (sometimes simply called the aperture) of the transducer 10will be between 4 and 5 mm. The reduced element count advantageouslyreduces the wire count (compared to conventional TEE transducers), whichmakes it easier to fit all the required wires into the narrower housing.The kerf K (i.e., the spacing between the elements) is preferably assmall as practical (e.g., about 25-30 μm or less). Alternative preferredtransducers may have between about 24-48 elements, spaced at a pitchbetween about 100-150 μm.

A second preferred transducer 10′ is shown in FIGS. 8A-8B. Thistransducer 10′ is similar to the first preferred transducer 10 describedabove in connection with FIGS. 7A-7C, except it is taller in theelevation direction. Similar reference numbers are used in both sets offigures to refer to corresponding features for both transducers.Numerically, the second transducer differs from conventional transducersin the following ways:

TABLE 2 conventional TEE second preferred Feature transducer transducerSize in the transverse (azimuthal) 10-15 mm about 4-5 mm directionNumber of elements 64 about 32-40 Size in the elevation direction 2 mmabout 8-10 mm Front face aspect ratio about 1:5 about 2:1(elevation:transverse) Operating frequency 5 MHz about 6-7.2 MHz

In alternative embodiments, the transducer 10 may be built with a sizein the elevation direction that lies between the first and secondpreferred transducers. For example, it may have a size in the elevationdirection of about 7.5 mm, and a corresponding elevation:transverseaspect ratio of about 1.5:1.

The transducer 10 preferably has the same transverse orientation (withrespect to the axis of the housing 60) as conventional TEE transducers.When the transducer is positioned in the stomach (as shown in FIG. 4),the image plane (azimuthal/radial plane) generated by the transducerintersects the heart in the conventional short axis cross-section),providing the trans-gastric short axis view of the heart, as shown inFIGS. 3 and 5. The transducer is preferably as wide as possible in thetransverse direction within the confines of the housing. Referring nowto the top view 22 in FIG. 7A, two examples of transducers that will fitwithin a 5 mm housing are provided in the following table, along with athird example that fits in a housing that is slightly larger than 5 mm:

TABLE 3 first second third Parameter example example example number ofelements in the transducer 38 36 40 a (azimuthal aperture) 4.75 mm 4.50mm 5.00 mm b (thickness) 1.25 mm 2.00 mm 2.00 mm c (inner diameter ofhousing at 4.91 mm 4.92 mm 5.39 mm the transducer) housing wallthickness 0.04 mm 0.04 mm 0.04 mm outer diameter of housing 4.99 mm 5.00mm 5.47 mmReferring now to the top view 22 in FIG. 8A, the three examples in Table3 are also applicable for fitting the second preferred transducer 10′within a 5-5.5 mm housing.

The above-describe embodiments assume that the housing is round.However, other shaped housings may also be used to house the transducer,including but not limited to ellipses, ovals, etc. In such cases,references to the diameter of the housing, as used herein, would referto the diameter of the smallest circle that can circumscribe thehousing. To account for such variations in shape, the housing may bespecified by its outer perimeter. For example, a 5 mm round housingwould have a perimeter of 5π mm (i.e., about 16 mm). When a rectangulartransducer is involved, using an oval or elliptical housing can reducethe outer perimeter of the housing as compared to a round housing. Forexample, an oval that is bounded by a 6 mm×2 mm rectangle with itscorners rounded to a radius of 0.5 mm contains a 5 mm×2 mm rectangularregion, which can hold the third example transducer in Table 3. Allowingfor a 0.04 mm housing wall thickness yields an outer perimeter of 15.4mm, which is the same outer perimeter as a 4.9 mm diameter circle. Thefollowing table gives the outer perimeters that correspond to some ofthe diameters discussed herein:

TABLE 4 outer diameter outer perimeter 2.5 mm 8 mm 4 13 5 16 6 19 7.5 24

Since the characteristics of the last one or two elements at each end ofthe transducer may differ from the characteristics of the remainingelements (due to differences in their surroundings), the last twoelements on each side may be “dummy” elements. In such a case, thenumber of active elements that are driven and used to receive would bethe total number of element (shown in Table 3) minus four. Optionally,the wires to these dummy elements may be omitted, since no signals needto travel to or from the dummy elements. Alternatively, the wires to maybe included and the last two elements may be driven, with the receivegain for those elements severely apodized to compensate in part fortheir position at the ends of the transducer.

Preferably, conventional beam-forming techniques are used to generateand aim a beam of acoustic energy in the desired directions. Forexample, focusing in the azimuthal direction may be accomplished byphasing (i.e., timing the excitation of individual elements L₁ . . .L_(N) in the array, and using appropriate time delays in the returns ofindividual elements before summing the respective returns into anultrasound return signal). Focusing in the elevation direction may beaccomplished based on the near-field and far-field properties of thesound signal, and will depend upon the physical height of the elementsin the elevation direction and optional acoustic lenses.

Resolution adequate to determine LV size and function depends upon acombination of resolution in azimuth, elevation, and axis. Thiscombination is referred to as “spatial resolution” and is illustrated inFIG. 9. FIG. 9 shows the image plane 320 and a scan line 310 that lieson the image plane 320. The axial direction AX is defined by the scanlines 310, and the transducer (not shown) is located far back along theAX axis. Out at the voxels being imaged, the azimuthal direction AZ isperpendicular to the AX axis within the image plane 320, and theelevation axis EL is perpendicular to the image plane 320. In an idealsystem, each voxel would be a point. In real-world systems, however, thevoxels have a volume that is defined by the resolution in all threedirections AX, AZ and EL, as shown for voxel 330. Similarly, while theimage plane 320 is depicted as a thin plane, the real-world image planewill have a thickness in the elevation direction EL that is equal to thethickness of the voxel 330 in the elevation direction.

The general formula for azimuthal and elevation resolutions is:

Δθ≈1.22λ/d,

where Δθ denotes the beamwidth in radians, λ the wavelength(corresponding to the transducer center frequency) and d the aperture inthe given direction (azimuth or elevation). The wavelength λ andaperture d are measured in the same units (e.g., μm).

Axial resolution depends indirectly upon the wavelength λ. Although theinventors are not aware of any specific formula for axial resolution, itis typically on the order of 16-64 times the wavelength. Thus,increasing the center frequency increases all three components ofspatial resolution. A center frequency on the order of 5-10 MHz is highenough to provide adequate resolution.

FIG. 10 illustrates the interplay between the three components indetermining the interaction between the shape of the resolution voxeland the boundary orientation in detecting and determining boundaries. Itshows the same voxel 330 that appears in FIG. 9, and also shows anillustrative piece 340 of the boundary being imaged that coincides withthat voxel. If the boundary orientation is random with respect to theresolution voxel, one suitable approach is to make the resolution voxelas cubical as possible. In order to obtain that shape, the azimuth andelevation resolutions for a given voxel should be approximately equal,which occurs when the front face of the transducer is approximatelysquare, as it is for the first preferred transducer discussed above inconnection with FIGS. 7A-7C.

For the first preferred transducer, the elevation aperture isapproximately the same as the azimuth aperture. In other words, thefront face of the transducer has a elevation:transverse aspect ratiothat is approximately 1:1 (i.e., it is approximately square). A squaretransducer with a width of 4-5 mm in the transverse direction wouldtherefore have an area of approximately 16-25 mm².

The formulas for azimuthal and elevation resolution are:

Δθ_(AZ)=1.22×λ/d _(AZ)

and

Δθ_(EL)=1.22×λ/d _(EL)

where Δθ_(AZ) and Δθ_(EL) are the azimuth and elevation resolutions,respectively, (both measured in radians); and d_(AZ) and d_(EL) are theazimuth and elevation apertures, respectively. These components may becombined into a single equation for overall resolution as a function ofarea and frequency, as follows:

Δθ_(OVERALL)=1.5×λ²/(d _(AZ) ×d _(EL))

As explained above, increasing the center frequency results in increasedresolution. However, increasing the center frequency also reduces thepenetration depth due to frequency-dependent attenuation, which isgoverned by the approximate formula

a≈0.5f×r

where a denotes the one-way attenuation in dB, f is the center frequencyin MHz, and r the depth in cm. Thus, one-way frequency dependentattenuation will typically be about 0.5 dB MHz⁻¹ cm⁻¹ and typicalround-trip frequency dependent attenuation will typically be about 1 dBMHz⁻¹ cm⁻¹.

The inventors have determined that a transducer center frequency betweenabout 6 and 7.2 MHz provides a good trade-off between resolution anddepth of penetration for TEE using a transducer with a 4.75 mm azimuthalaperture. In the embodiments described herein, that range of frequenciescan typically provide enough depth of penetration to image the far wallof the left ventricle (in the TGSAV) so that the interior volume of theleft ventricle can be computed. (In most subjects, a 12 cm depth ofpenetration is adequate to image the far wall. For many subjects, adepth of penetration of about 9-10 cm will suffice).

When the transducer elements are spaced at a 125 μm pitch, using atransducer center frequency of 6.16 MHz is particularly advantageousbecause it corresponds to a wavelength of λ=250 μm. At that wavelength,the elements are spaced at a pitch of 0.5λ, which is sometimes referredto as “half-wavelength pitch”. As is well known to those skilled in theart, a half-wavelength pitch is excellent for eliminating grating lobeswhile still minimizing element count for a given azimuth aperture.Somewhat larger pitches, e.g. 0.6λ, still work reasonably well in termsof eliminating grating lobes. Thus, for transducers that can operate ata range of center frequencies, acceptable performance can be maintainedeven if the frequency is increased about 20% (i.e., to the point wherethe pitch becomes about 0.6λ).

As explained above, the formula for the angular resolution is θ≈λ/d.Referring to the tables above, the first example of the first preferredtransducer has a 38 element transducer with a 125 μm pitch, resulting ina 4.75 mm transducer width (d=4750 μm). It is preferably approximatelysquare and operates at a center frequency of 6.16 MHz (λ=250 μm). Whenthose values for d and λ are plugged into the equation for resolution,the result is 0≈0.053 radians, which converts to approximately 3 degreesresolution in both azimuth and elevation.

The increased size of the transducer in the elevation direction helpsimprove the angular resolution of the system in the elevation direction(as compared to a conventional TEE transducer with a 2 mm elevation).This increased resolution in the elevation direction helps compensatefor losses in angular resolution in the azimuth direction caused byshrinking the azimuthal aperture down to about 4-5 mm.

The inventors have noticed that increasing the size of the transducer inthe elevation direction further, so that it is larger than the size inthe azimuthal direction provides improved performance when imaging thefar wall of the heart in the TGSAV. This increase in transducerelevation causes the resolution voxel to shrink in the elevationdirection at distances that correspond to the far wall of the LV, whichresults in increased resolution in the elevation direction. Theinventors believe that increasing resolution in this direction ishelpful at least in part because the far wall is slanted about the Yaxis with respect to the front face of the transducers. (The Y axis isshown in FIG. 8B.) Shrinking the size of the voxel in the elevationdirection therefore minimizes the variations of the components of returnsignals arising from specular reflections that fall within a singlevoxel.

The inventors have determined that the images of the TGSAV are betterwhen the transducer is more than 1.5 times as large in the elevationdirection as the transverse direction, and that the best images of theTGSAV are obtained when the transducer is about two times as large inthe elevation direction as the transverse direction, as it is for thesecond preferred transducer 10′ described above in connection with FIGS.8A and 8B and Table 2.

Instead of the 90° sector width that is typically used in conventionalTEE systems, the preferred embodiment uses a smaller sector width (e.g.,60 degrees). Referring now to FIG. 11, a 60° sector 92 is shownemanating from the front face 14 of the transducer 10. The effectiveazimuthal aperture at an angle θ from the centerline CL can be obtainedby multiplying the (nominal) azimuthal aperture (at θ=0) by cos(θ).Since)cos(30°=0.866 and)cos(45°=0.707, restricting the sector width to60° (i.e., 30° on each side of the center line CL) causes a smallerdegradation in worst-case azimuthal aperture: azimuthal aperture isdegraded by only 13.4%, compared with 26.8% in the case of a 90° sectorwidth. For example, the worst case aperture for a 4.75 mm widetransducer (5 mm housing diameter) in a 60° sector would be about 4.11mm. The result is improved effective azimuthal aperture, which improvesthe overall resolution obtainable with small transducers. If aconventional 90° sector were to be used, a 5.82 mm wide transducer (6.1mm housing diameter) would be needed in order to provide the same worstcase aperture.

After a beam of ultrasound energy is sent into the patient using thetransducer described above, the ultrasound return signal is received,preferably by the same transducer. The transducer converts theultrasound return signal into an electrical return signal. This processcontinues as the beam is swept through the imaging sector. FIG. 12 is aschematic illustration of the path of the ultrasound beam as it is sweptthrough the sector, first along line B₁, then along line B₂, andcontinuing on through line B_(M). These scan lines B₁ . . . B_(M)correspond to the fan-shaped beam 90 (shown in FIG. 4) and the sector 92(shown in FIG. 11). Although the illustration only includes a smallnumber (M) of scan lines, an actual system would have many more scanlines that are much more densely packed, so as not to adversely impactthe azimuthal resolution.

The electrical return signal can be modeled as being anamplitude-modulated signal, with the carrier frequency at the centerfrequency, and with the modulation being caused in large part fromscatterer spacing and other tissue characteristics such as the presenceof connective tissue around heart muscle bundles. The electrical returnsignal is demodulated and digitized (i.e., sampled) to form ademodulated and digitized return signal (DDRS). A variety ofconventional techniques that are well known to persons skilled in therelevant arts may be used to form the DDRS. One example is to digitizethe electrical return signal and then rectify the result (i.e., take theabsolute value) to form a rectified digitized ultrasound return signal.Another example is to rectify the electrical return signal in analogform, and then digitize the result to form the DDRS. Alternativedemodulation approaches may also be used to extract the modulationinformation from the electrical return signal, including but not limitedto coherent demodulation, Hilbert transforms, and other demodulationtechniques that are well known to persons skilled in the relevant arts.

FIG. 13 is a schematic illustration of the DDRS that corresponds to asection of one of the ultrasound beams B₁ . . . B_(M) of FIG. 12. Eachsample is represented by a dot S0 . . . S143. Each sample corresponds toa point in 2D space based on the direction of the beam and the time ittook for the signal to travel from the transducer to the point inquestion and back. For example, if the return signal is digitized at 50MHz, the time between samples will be 0.02 μs, which corresponds to adistance of 0.015 mm (based on the speed of sound in the body). Althoughthe illustration only includes only 144 samples, an actual system wouldhave many more samples in each scan line to provide the desiredresolution. For example, to obtain a depth of penetration of 12 cm withthe samples spaced 0.015 mm apart, 8000 samples would be needed. Becausethe beam of ultrasound energy is swept about a center point, polarcoordinates are useful to organize the samples, at least in this stageof the processing. In some embodiments, the samples are analyzedentirely in polar coordinates, and only converted into rectangularcoordinates for viewing on a conventional computer monitor. In otherembodiments, the sample space may be converted into rectangularcoordinates at an earlier stage of processing. The remaining explanationconsiders coordinates along each scan line (constant θ in the (r, θ)polar coordinate system, with r varying along the scan line), and pixeldata associated with the center of the pixel along the scan line.Conversion to a sector image is well known in the field of ultrasoundimaging.

The samples of each scan line are preferably processed by two differentalgorithms: one algorithm that analyzes intensity characteristics of thesamples, and one algorithm that analyzes frequency characteristics ofthe samples.

For the first algorithm (i.e., the intensity algorithm) the samples ofthe scan line is divided into a plurality of pixels, with each pixelcontaining a plurality of samples. In the FIG. 13 example, each pixelcontains 16 samples, as indicated by the boxes labeled “WIAP j” (whichstands for “Window for Intensity Algorithm for Pixel j, where j is aninteger from 0-8) that appear below the corresponding samples. Pixeldata generated by signal processing is associated with the centerposition of the corresponding pixel. Of course, other numbers of samplesper pixels could also be used instead of 16. In one preferredembodiment, for example, each pixel contains eight samples. Theintensity algorithm is preferably a conventional image processingalgorithm that converts the samples into a conventional image. Theintensity for any given pixel is determined based on the amplitude ofthe samples that correspond to that pixel, with higher intensitiescorresponding to larger amplitudes. In the case of a 16 sample pixel,the average of those 16 samples would be used to determine the intensityat the pixel (with higher average intensity values appearing brighterand lower average intensity values appearing darker). Optionally, theintensity level of the pixel (or the samples that make up that pixel)may be compressed using conventional procedures such as logarithmiccompression.

The second algorithm (i.e., the frequency algorithm) analyzes thefrequency characteristics of the sample space and determines the spatialfrequencies in scatterer spacing. Examples of suitable algorithms aredescribed in U.S. Pat. No. 5,417,215 (hereinafter “the '215 patent”),which is incorporated herein by reference. The article “SpectralAnalysis of Demodulated Ultrasound Returns: Detection of ScattererPeriodicity and Application to Tissue Classification” by S. Roth, H. M.Hastings, et al., published in Ultrasonic Imaging 19 (1997) at pp.266-277, is also incorporated herein by reference.

The frequency algorithm provides a second result for each pixel in theimage (i.e., in addition to the result produced by the intensityalgorithm). Because most frequency analyzing algorithms provides betterresults when a larger number of data samples is used, and because eachpixel only has a limited number of samples, samples on either side ofthe pixel in question are preferably combined with the samples in thepixel itself to increase the number of samples. In the illustratedexample, each pixel contains 16 samples, but the frequency algorithm forany given pixel operates on 64 samples that are preferably centered inthe pixel, as indicated by the boxes labeled “Window for Freq. Algorithmfor Pixel k” (where k is an integer from 2-6) that appear below thesamples. In this case, for example, all the samples from pixels 2-4 plushalf the samples from pixels 1 and 5 would be used to perform thefrequency analysis for pixel 3. Of course, other numbers of samplescould be used for the frequency analysis instead of 64. Powers of 2,however, are preferable when a fast Fourier transform (FFT) algorithm isused. Optionally, windowing techniques (such as Hamming windows) may beused to weight the samples in the center more heavily than the samplesthat are near the ends.

FIG. 14 is a flowchart of a suitable frequency algorithm. In thisalgorithm, steps 1 and 2, taken together, attempt to discern thematerial that the pixel in question is made of (and more specifically,whether that pixel is blood or muscle) based on the frequencycharacteristics of samples in the pixel and the samples in theneighboring pixels.

In step 1, a Fourier analysis is performed on the samples to determinethe power distribution in the various frequency bands at each pixel. Theend result of the Fourier analysis of step 1 is a set of amplitudecoefficients for each of a plurality of different frequencies, for eachof the pixels (i.e., one set of coefficients for the first pixel, asecond set of coefficients for the second pixel, etc.). The Fourieranalysis may be implemented using any of a variety of algorithms thatare well known to persons skilled in the relevant arts (e.g., aconventional FFT algorithm). Alternative embodiments may use otherfrequency analysis tools to achieve similar results, such as bandpasstechniques (preferably integer-based FIR recursive), wavelet techniques,etc. In step 2, the ratio of power in a selected frequency band to thepower in the entire spectrum for each pixel is computed. Thus, for eachpixel, the following formula applies:

R=E _(BAND) /E _(TOTAL)

Where E_(BAND) is the power in the selected frequency band, E_(TOTAL) isthe total power in the portion of the spectrum, and R is the ratio ofthose two powers. When a Fourier analysis is used, the power in anygiven band equals the sum of the squares of the amplitudes of theFourier coefficients within the band. The “selected frequency band” inthis step is preferably selected so that changes in the ratio R arecorrelated to differences in the material that is being imaged (e.g.,blood v. muscle). Alternatively, it may be selected so that changes inthe ratio R are correlated to differences in S/N ratio, with larger Rsbeing correlated with signal and smaller Rs being correlated withspeckle or electric noise. Optionally, different “selected frequencybands” may be used for near returns and for distant returns. Forexample, a wider frequency band may be used for signals that correspondto distant structures. In other words, the band selection can be afunction of depth.

One suitable set of numeric values that results in a correlation betweenR and the material being imaged will now be discussed. Consider firstthe ultrasound return from a single scatter at a depth of r mm. Thereturn from this scatter arrives after a time delay of t μs, given by

t=r/v=r/(0.77 mm/μs)=1.30 r μs,

where the scaling factor of 0.77 mm/μs represents a round trip from thetransducer to the scatterer and back (assuming the velocity of sound intissue is to be 1.54 mm/μs).

The effects of scatterer periodicity upon the spectrum of thedemodulated ultrasound return may be calculated in the case ofseparations large enough so that the ultrasound returns do not overlap(i.e., separations Δr larger than Δr₀=0.77 mm/μs×Δt). For example, inthe case of an ideal one cycle pulse, a 5 MHz center frequency, and aideal wide-bandwidth transducer,

Δt=1/f _(c)=1/(5 MHz)=0.200 μs,

and thus

Δr ₀=0.77 mm/μs×0.200 μs=0.154 mm.

The internal structure of cardiac muscle displays variations on this andlarger spatial scales. In contrast, scattering from blood ischaracterized by full-developed speckle, including variations on all,and especially much smaller spatial scales. As a result, low frequenciesare indicative of muscle, and high frequencies are indicative of blood.This suggests defining the upper limit of a low frequency band to beless than about 4 MHz, corresponding to a minimal spatial scale Δr_(MIN)of

Δr _(MIN)=0.77 mm/μs×1/(4 MHz)=0.77 mm/μs×250 μs=0.193 mm.

The inventors have performed tissue experiments that used a signaldigitized at 50 MHz (corresponding to sampling interval of 0.02 μs), andcomputed the FFT in a 64 point window (corresponding to 64×0.02 μs=1.28μs, or 0.986 mm). With that size window, the inventors selected a lowfrequency band that included Fourier frequencies of between 2 and 5cycles per window (inclusive), which corresponds to frequencies between2/1.28 MHz=1.56 MHz and 5/1.28 MHz=3.91 MHz.

Using the formula for R set forth above (R=E_(BAND)/E_(TOTAL)) for thislow frequency band, the ratio of Fourier power in the low frequency bandto the total Fourier power is computed for each pixel. The end result ofstep 2 in FIG. 14 is a value of R for each pixel.

The inventors have found that, for the parameter values used in thisexample, R-values of around 0.45 are significantly correlated to thepresence of muscle tissue at the pixel of interest, and R-values ofaround 0.20 are significantly correlated to blood or regions dominatedby electronic noise. The remaining part of the algorithm uses thisinformation to improve the image by increasing the intensity of theportions of the image that correspond to muscle and decreasing theintensity of the portions of the image that correspond to blood. Sinceblood is less reflective than muscle, this difference enhances thecontrast between blood and muscle.

The inventors have determined that cardiac ultrasound images aredramatically improved when the intensity of the areas with R-valuescorresponding to muscle is increased to about 120% of its originalvalue, and when the intensity of the areas with R-values correspondingto blood is decreased to between about 20% and 50% of its originalvalue. Thus, in step 3 of FIG. 14, a gain factor of about 1.2 isassigned to those portions of the image with R values of about 0.45, anda gain factor of between about 0.2 and 0.5 is assigned to those portionsof the image with R values of about 0.20. This gain factor is referredto herein and a “feature gain factor” or FGF because the gain is featuredependant.

While most pixels in most images will have R-values that permit thepixel to be classified as being either muscle or blood, in some casesthe classification is less clear. For example, pixels that straddle aboundary between muscle and blood have less predictable R values. Inaddition, although the R-values from blood may average out to 0.20, anygiven pixel of blood may vary widely from that R-value due to randomstatistical variations. Accordingly, a monotonic, preferably smoothfunction may be used to map R to FGF in some embodiments. FIG. 15 is anexample of a suitable function for this purpose. Optionally, additionalrestrictions may be built into the mapping function, based on othertissue characteristics.

Finally, in step 4 of FIG. 14, the results of the intensity algorithmand the frequency algorithm are combined by multiplying the intensityvalue for each pixel (obtained from the intensity algorithm) by the FGFvalue for that pixel (obtained from the frequency algorithm). The resultis an enhanced image in which the pixels that are probably muscle havebeen brightened while the pixels that are probably blood have beendimmed. This enhanced image is then displayed using conventionalhardware and software techniques (including, for example, usinginterpolation to convert the polar coordinates to rectangularcoordinates).

The actual choice of the Fourier frequency bands, R-values andcorresponding FGF values depends upon a variety of factors including butnot limited to transducer center frequency, sampling rate, window sizeand any optional windowing techniques used in signal processing,transducer bandwidth, the width of interrogating pulse, etc. In oneembodiment, for example, a transducer center frequency of 7.5 MHz isused, the scan line is digitized at about four times the centerfrequency (i.e., about 30 MHz), and the distance between the samples isabout 0.026 mm.

In alternative embodiments, other normalized (i.e.,non-amplitude-dependent measures) may be used instead of dividingE_(BAND) by E_(TOTAL). For example, the ratio of power in a firstfrequency band to the power in a second frequency band may be used tocompute R, as explained in the '215 patent (e.g., by dividing E_(BAND1)by E_(BAND2)). In alternative embodiments, two or more Fourier analysesmay be performed for each pixel, using a corresponding number of linesof samples, where the center of each line is contained within the pixel.For example, a two line per pixel arrangement in which a first 1DFourier analyses is implemented along a line of samples in the radialdirection, and a second 1D Fourier analyses is implemented along asecond line of samples in the tangential direction. The results fromthose two lines of samples are then merged (e.g., by averaging). Instill other embodiments, a 2D Fourier algorithm may be used instead ofthe 1D algorithms described above.

Ordinarily, the above-described operations are performed on theuncompressed image data. Under certain circumstances, however, it may bepossible to perform corresponding operations directly on a compressedversion of the image data.

Once the enhanced images have been generated, they may be displayedusing conventional hardware. The images may be continuously updated anddisplayed for the entire time that the probe is in position, so that thephysician can visualize the patient's heart in real time. In alternativeembodiments, images may be acquired and optionally stored periodically(e.g., by capturing one or more complete heartbeats every two minutes).Optionally, the ability to compare a prior heartbeat to the currentheartbeat may be provided by, for example, playing back a stored videoclip (or “loop”) of an old heartbeat in one window, and displaying thecurrent image in a second window.

In contrast to conventional extended duration TEE using a transducerwith a 10-15 mm azimuthal aperture, which is ordinarily done only undergeneral anesthesia in the closely monitored environment of an room, thesmaller diameter of the preferred embodiments described herein permitsthe preferred embodiments to be used without general anesthesia, and inless closely monitored environments. Optionally, the preferredembodiments may be used with sedation or local anesthesia in place ofthe general anesthesia that was used with conventional extended durationTEE. It may even be possible to forgo the use of sedation or anesthesiaaltogether. In such cases, the patient may optionally be medicated withan analgesic.

Optionally, regions of high relevance as detected by the feature gainfactor may be highlighted, typically using colorization, whilepreserving the intensity of the gray-scale image, as explained in the'215 patent. Additional techniques for image enhancing can be found inapplication Ser. No. 10/633,949, filed Aug. 4, 2003, and entitled“Method and Apparatus for Ultrasonic Imaging,” which is incorporatedherein by reference.

The preferred embodiments described above advantageously permitnon-invasive, intermediate and long-term monitoring of cardiac functionusing a small transducer that fits into a housing approximately 5 mm indiameter, thereby reducing or eliminating the need for anesthesia. Thepreferred embodiments described above combine a plurality of techniquesto produce images that are comparable to or better than images that wereconventionally obtained by much larger transducers. The images producedby the preferred embodiments described above are repeatably and reliablyusable for monitoring heart function, with adequate penetration depth tosee the far wall of the left ventricle (10-12 cm) and adequateresolution to determine LV size and function from an image of theendocardial wall in real time, despite the use of a smaller transducer.Thus, in contrast to prior art systems which provide a depth ofpenetration that is less than 15 times the azimuthal aperture of thetransducer (e.g., obtaining 10 cm penetration using a 10 mm transducer)the preferred embodiments can provide penetration that is greater than15 times the azimuthal aperture of the transducer, or even greater than20 times the azimuthal aperture of the transducer (e.g., obtaining 10 cmpenetration using a 4.75 mm transducer).

The preferred embodiments described above use a probe that is muchnarrower than conventional TEE probes, and may be used to monitor heartfunction over an extended period of time and to obtain an understandingof the patients' hemodynamic status. Such information may be useful inchoosing treatments and improving outcome in many situations (includingbut not limited to critical medical problems such as hypotension,pulmonary edema and heart failure).

The above-described embodiments permit direct visualization of cardiacfunction, which permits evaluation of a patient's hemodynamic statusincluding intravascular volume (normal, low or high), cardiaccontractility (how well the left ventricle pumps), cardiac ischemia(inadequacy of blood flow to the heart muscle) and cardiac tamponade(fluid in the pericardial sac limiting heart function). For example,information about intravascular volume status can be derived fromdirectly visualizing the size of the left ventricle and monitoringchanges in size with treatments over time. Information aboutcontractility can be obtained by directly visualizing the contraction(pumping) of the left ventricle, either using qualitative visualestimates or quantitatively. Information about ischemia is availableduring direct visualization of the left ventricle since ischemia resultsin abnormal motion of the walls of the left ventricle (wall motionabnormality). Information about possible cardiac tamponade orpericardial effusion (fluid in the heart sac) is available when usingultrasound to directly visualize the heart.

The narrowness of the probe may enable the above-described embodimentsto provide this information for longer periods of time, outside theoperating room, and/or without anesthesia. The above-describedembodiments also lend themselves to use in settings where interventionalcardiac procedures are performed such as the cardiac catheterization andelectrophysiology laboratories, both for monitoring the effects ofphysicians' interventions on cardiac and hemodynamic function and forguiding the placement of devices. For example, they may be used to helpthe physician correctly place the pacing leads to achieve the desiredresult. The above-described embodiments may also be used in non-cardiacapplications in which a narrower probe is needed or beneficial.

The above-described embodiments are not limited to ultrasound imagingmodes, and may be used in alternative ultrasound modes (e.g., pulsedwave Doppler, continuous wave Doppler, and color flow imaging Dopplermodes). These alternative modes may be performed using the sametransducer as the above-described imaging modes and may yieldinformation which can be combined with images, optionally in real-time.For example, color flow Doppler information may be obtained duringimaging of the mitral valve (between the left atrium and left ventricle)while maintaining transducer position in the mid to lower esophagus.Such an application would permit evaluation of leakage of the mitralvalve (mitral regurgitation or insufficiency).

If desired, the preferred embodiments described above may be scaled downfor neonatal or pediatric use. In such cases, a transducer that isbetween about 2.5 and 4 mm in the azimuthal direction is preferable,with the elevation dimension scaled down proportionally. Because lessdepth of penetration is required for neonatal and pediatric patients,the operating frequency may be increased. This makes λ smaller, whichpermits the use of a smaller transducer element spacing (pitch), and acorrespondingly larger number of elements per mm in the transducer. Whensuch a transducer is combined with the above-described techniques, theperformance should meet or surpass the performance of conventional 7.5mm TEE probes for neonatal and pediatric uses.

The embodiments described herein may also be used in non-cardiacapplications. For example, the probe could be inserted into theesophagus to monitor the esophagus itself, lymph nodes, lungs, theaorta, or other anatomy of the patient. Alternatively, the probe couldbe inserted into another orifice (or even an incision) to monitor otherportions of a patient's anatomy.

If desired, the center frequency may be lowered (e.g., down to about 4.5MHz) to provide additional depth of penetration when needed (e.g., forvery large patients). Although this will also reduce the resolution, theresult may be acceptable when very large structures are being imaged.Alternatively, the transducer size and housing diameter may be scaled upin size (e.g., to about 7 mm) if the reduced resolution results inunusable images.

Numerous alternative and optional features may be substituted and addedto the above-described embodiment. One optional feature is digitalbeamforming using significant oversampling. For example, if thetransducer is operated at 7 MHz, and the return is digitized at30×frequency, 30×7 MHz=210 MHz digitization would be required. That datacould then be downsampled by a factor of five to reduce the number ofdata points to a 42 MHz sample. Such downsampling would reduce the noisefloor due to front-end noise by a factor of ≈5, (i.e., over 2 bits inpower). Similarly, downsampling by a factor of 7 would reduce the noisefloor by a factor of 17.

FIG. 16A depicts the front face of an alternative 2D transducer 500,which includes a 2D array of active elements 510. The concepts describedherein can also be implemented using this type of transducer by makingappropriate adjustments that will be apparent to persons skilled in therelevant arts.

FIG. 16B depicts the front face of another alternative 2D transducerdesign that is referred to as a “sparse 2D transducer.” The sparse 2Dtransducer 600 has a column 610 of “transmit” elements 611, used fortransmitting the ultrasound, and a row 620 of receive elements 621 usedfor receiving the ultrasound signal. As shown, there is one element 630common to both the column 610 of transmit elements and the row 620 ofreceive elements. This common element 630 may be used for transmission,reception, or both. This transducer design reduces electronic noise byusing separate transmit and receive elements, which eliminates the needfor electronic transmit/receive switches at the elements. The conceptsdescribed herein can also be implemented using this type of transducerby making appropriate adjustments that will be apparent to personsskilled in the relevant arts.

Alternative embodiments of the invention may use fewer techniques and/orimplement those techniques to a lesser extent, and still maintain theability to produce an acceptable image. For example, depending on theother components in the system, it may be possible to obtain anacceptable image using a 75° sector width, or even using a 90° sectorwidth. It may also be possible to obtain an acceptable image using atransducer with an elevation:transverse aspect ratio of about 2:3 inplace of the preferred 1:1 or 2:1 aspect ratios. Another alternativewould be to use some or all of the above-described techniques with atransducer that is slightly larger than the preferred embodimentsdescribed above, yet still smaller than conventional 10 mm TEEtransducer. Numerous other modifications to the above-describedembodiments will be apparent to those skilled in the art, and are alsoincluded within the purview of the invention.

We claim:
 1. A system for imaging regions that include at least twotypes of tissue, the system comprising: an ultrasound imaging system;and a probe including (a) a housing having a distal end and a flexibleshaft, (b) an phased array ultrasound transducer mounted in the distalend of the housing, the transducer being made of a stack of piezoelement spaced at a pitch between 125 and 128 μm with a kerf of lessthan 30 μm, and (c) an interface that operatively connects theultrasound transducer to the ultrasound imaging system such that theultrasound imaging system drives the ultrasound transducer and receivesreturn signals from the ultrasound transducer, wherein the distal endhas an outer diameter of less than 7.5 mm and the flexible shaft has anouter diameter of less than 7.5 mm, wherein the transducer is operatedat between 6 and 7.2 MHz, and wherein the ultrasound transducer isconfigured to emit a fan-shaped beam with a sector width of 60° or less.2. The system of claim 1, wherein the distal end has an outer diameterof less than 6 mm.
 3. The system of claim 1, wherein the distal end hasan outer diameter of about 5 mm.
 4. The system of claim 1, wherein thetransducer is operated at 6.16 MHz.